The following information is provided to assist the reader to understand the technology described below and certain environments in which such technology can be used. The terms used herein are not intended to be limited to any particular narrow interpretation unless clearly stated otherwise in this document. References set forth herein may facilitate understanding of the technology or the background thereof. The disclosure of all references cited herein are incorporated by reference.
Heart failure, or the inability of the heart to pump sufficient blood for the body's needs, results in very poor quality of life, huge costs to society, and hundreds of thousands of yearly deaths. Heart failure is caused by an abnormally low cardiac output. Cardiac output is the outflow of blood from the heart and can be measured in liters of blood flow per minute or LPM. Cardiac output for a normal man at rest or during light activity is approximately 5 liters per minute. Severe heart failure exists when the cardiac output is between approximately 2.5 to 3.5 liters per minute. For an average man in heart failure having a heart rate of 80 beats per minute, the average amount of blood that is pumped with each heartbeat (sometimes referred to as stroke volume) might, for example, be 37 milliliters or ml. If the same man was not in heart failure, his heart might, for example, pump 62 milliliters with each heartbeat. An effective treatment for such heart failure would be to increase the low, 37 ml stroke volume up to the normal, 62 ml stroke volume.
The main pumping chamber of the heart or left ventricle or LV, which has an inlet mitral valve and an outlet aortic valve. During left ventricular contraction or systole, the inlet valve closes as blood is pushed through the aortic valve into the aorta or main artery to the body. When the LV is resting during diastole, LV pressure may be between 2 and 20 mm of Hg pressure. This diastolic pressure is termed the LV preload. The preload will be in the higher end of its pressure range during heart failure. During active LV contraction or systole, the LV must eject its blood against the pressure in the aorta. Aortic pressure is typically between 70 and 140 mm Hg Pressure. This aortic pressure is termed the after-load. It is well known that, if the after-load is reduced in heart failure, the LV stroke volume will naturally increase and this increase is one reason that afterload-reducing drugs such as ACE-inhibitors help heart failure patients.
Heart failure is generally caused by one of two disease mechanisms. The first is a physical weakness of LV muscular contraction, typically as a result of LV muscle damage from one or more heart attacks. Heart failure resulting from a weak heart muscle combined with more than adequate blood volume stored in the LV at the beginning of the systolic muscular contraction for producing a normal stroke volume is termed “systolic failure”. In systolic failure, the ratio or fraction of the ejected stroke volume compared to the starting LV blood volume is low. This fraction is known as the ejection fraction or EF. Normal EFs are in the range of 55 to 65 percent, but with systolic heart failure, EFs are typically in the 10 to 40 percent range.
A second form of heart failure is known as “heart failure having preserved ejection fraction”, or by the somewhat outdated term, “diastolic failure”. In diastolic failure, the heart muscle is abnormally stiff and the heart is not able to adequately fill with blood. The stroke volumes in heart failure having preserved ejection fraction are lower than normal, but the EF is normal or near normal.
Blood pumps which lower the aortic pressure after-load can be desirable because such pumps allow the failing LV to eject more blood with less effort. However, no commercially available afterload reducing devices have thus far been shown to be practical for extended support of the failing LV. Instead, all long term (that is, months to years), commercially available heart assist devices, whether rotary turbine pumps or collapsing chamber pumps, go around or bypass the failing LV, pumping blood from the LV apex through the pump into the aorta. By doing so, those pumps act in parallel to the LV and compete with the LV in their pumping action. This pumping competition has several negative complications including right heart failure, fusion of the aortic valve over time and the risk of collapsing the LV. Collapsing chamber pumps are physically large and thus cannot be implanted in some small patients. Rotary turbine pumps are desirably smaller, but have other limiting complications. For example, rotary turbine pumps induce high levels of shear stress in the blood elements and also may reduce the normal pulsatility of the blood entering the aorta. High shear stress on the blood cells promotes blood clotting which can lead to strokes and heart attacks. Physicians try to reduce this blood clotting by giving the patients anticoagulants, which, in turn, puts the patients at risk of excessive bleeding. These clotting and bleeding complications are substantial limitations to broader use of rotary turbine assist pumps.
A number of moving valve pumps have been disclosed for assisting blood flow. For example, U.S. Pat. Nos. 5,676,162, 5,676,651, and 5,722,930, disclose a single-stroke, moving valve pumps designed for ascending aortic placement. That device uses a commercially available artificial heart valve with attached magnets and requires excision of a portion of the aorta. A series of separate electric coils step the valve/magnet combination forward in a sliding action within a cylinder. The device is quite large for the limited space available between the heart and the take-off vessels from the aorta to the upper body and brain. The device is designed to have one stroke in synchronization with each LV systole. The blood volume required for closing commercially available heart valves is typically 2-5 ml and therefore multiple smaller oscillations per heart contraction in such devices would suffer from volumetric inefficiency. Another problem with such devices is the tight crevice between the cylinder wall and the moving valve. This tight space results in high blood shear and the corresponding risk of stroke or blood clotting complications if anti-coagulant therapy is necessary. The same problem exists with a moving valve pump disclosed in U.S. Pat. No. 4,210,409, which included two valves (one stationary and one moving).
U.S. Pat. No. 5,147,281 discloses an oscillatory valve blood pump that is external to the body and fits in an enclosure the size of a briefcase. The pump uses a stationary coil to attract a magnetic tube encasing a one-way valve. A forward stroke of the one-way valve propels blood until the tube assembly stops and is repelled backward by return leaf springs that are charged during the forward stroke. A second stationary valve is sometimes in the circuit. A stretchable silicone rubber tube connects the tube or pipe-valve assembly with the pump inlet and outlet.
Nitta, S. et al., “The Newly Designed Univalved Artificial Heart,” ASAIO Transactions Vo. 37, No. 3, M240-M241 (1991) describes a “univalved artificial heart” powered electro-magnetically wherein the valve oscillates within a frequency range of 1 to 30 Hz. The valve is contained in a tube, with attached magnetic material. Stationary electric coils actuate the tube-magnet-valve combination. The valve is described as a jellyfish valve. A problem with jellyfish valves is the compound curvature or wrinkling of the membrane that occurs when the valve opens and closes. One can liken the action of the jellyfish valve to that of an umbrella that oscillates between a circular flat membrane and a wrinkled umbrella shape as it closes and opens. Wrinkling of the membrane is virtually impossible to prevent in a jellyfish valve and introduces stresses and strains that significantly limit the life of the valve.
U.S. Pat. No. 5,266,012 also uses a jellyfish valve in a vibrating pipe blood pump intended for use outside the body. Because the vibrating tube pump portion is separable from the drive mechanism. the blood-contacting portion of the pump is disposable.
U.S. Pat. No. 7,588,530, describes a moving valve pump having a curved blood flow path as well as a moving valve pump having a linear blood flow path. U.S. Pat. No. 7,588,530 discloses various drive mechanism to oscillate the moving valve in synchronization with the R wave of the patient's electrocardiogram. In the case of a pump having a linear blood flow path, a linear motor is disclosed to drive the moving valve thereof. U.S. Pat. No. 7,588,530 further discloses moving valves including a plurality of openings or ports wherein each port includes a resilient flap of material to open the port upon rearward movement of the moving valve and close the port upon forward movement of the moving valve. U.S. Pat. No. 7,588,530, further discloses movement of the moving valve thereof in the latter half of systole.
Numerous pharmacologic, biologic, and mechanical interventions have been devised to address heart disease/failure. Nonetheless, heart failure remains a major public health problem with an estimated five million victims in the United States alone.